Intraocular lenses (IOLs) were first used as a replacement for damaged natural crystalline lenses in 1949. These early IOL experiments were conducted in England by Dr. Howard Ridley an RAF ophthalmologist. Dr Ridley first observed acrylate polymer biocompatibility in the eyes of pilots who had sustained ocular injuries from polymethylmethacrylate (PMMA) shards when their aircraft canopies were shattered.
However, it took nearly thirty years for ophthalmologists to embrace IOL implantation as a routine method for restoring vision in patients suffering from diseased or damaged natural crystalline lenses. Early IOLs were made from PMMA because of its proven biocompatibility; however, PMMA IOLs require a 5 mm to 7 mm incision. Incision size is directly related to patient trauma, discomfort and healing times. Moreover, incisions sizes in the 5 mm to 7 mm range generally require sutures further increasing procedural complexity and patent discomfort.
Lens size dictates incision size and lens size is in turn determined by the size of the capsular sac and natural crystalline lens. Thus lenses made from a rigid polymer such as PMMA require an incision size at least as large as the minimum IOL dimension which is generally 5.5 mm on average. In an effort to decrease incision size and corresponding patient discomfort, recovery time and procedural complexity a number of IOL designs suitable for insertion through small incisions have been developed; most notably foldable IOLs. Foldable IOLs are made from non-rigid, or flexible polymers including hydrophobic acrylics, hydrophilic hydrogels, silicone elastomers and porcine collagen. Intraocular lenses made form these materials can be folded or rolled into implantable configurations having minimum dimensions suited for 3 mm incisions, or less.
Intraocular lenses made from flexible polymers are not easily manipulated especially after being rolled or folded. Thus these lenses must be manipulated using devices specifically manufactured for inserting the lens into the capsular sac through minute incisions. The IOL must be tightly folded or rolled prior to placing it into the inserter tip, or barrel. The smaller the incision size, the more tightly the IOL must be folded or rolled. Devices designed specifically to insert an IOL into the capsular sac are referred to generally as inserters.
The typical inserter is similar to a syringe in that it comprises a plunger-like device that engages the folded or rolled IOL restrained within a barrel-like tip. As pressure is exerted on the plunger the IOL is pushed out of the tip and into the eye. Once inside the capsular sac the IOL unfolds. The IOL may also include haptics which are spring-like arms that help hold the IOL in place. Sutures are generally not required with modern IOLs.
The IOL inserter barrel is generally made form polymers such as polyolefins which are highly hydrophobic. When a polymer IOL is pushed through the polyolefin barrel frictional forces impede the IOL's progress requiring increasing amounts of force. As the pressure increased the folded polymer IOL will tend to expand circumferentially inside the inserter as longitudinal movement is restricted by friction. If the friction coefficient of the tip relative to the lens is too great the lens may seize in the inserter tip making IOL delivery impossible. Moreover, the inserter tip may crack (craze) or even fracture as longitudinal pressure is increased resulting in IOL delivery failure.
In an effort to minimize friction within the inserter tip and ease IOL deployment numerous lubricious coatings have been developed. The lubricious coatings are generally composed of biocompatible hydrophilic polymers applied directly to the inserter interior surface. However, as incision sizes become progressively smaller, the inserter size begins to reach a diameter whereby lubricants alone no longer prove the lubricity required to overcome frictional forces. Furthermore, the higher the coefficient of friction between the IOL and lubricated inserter barrel, the greater the possibility that lubricant will be stripped from the barrel's interior surface and transferred to the IOL. Therefore, there is a need for IOL materials having inherently lower coefficients of friction relative to the lubricated inserter barrel. Moreover, some patients may be so hypersensitive to available lubricants that lubricant-free IOL delivery may be required in extreme cases. In these cases it is essential the IOL material have the lowest possible coefficient of friction.
Therefore, it is an object of the present invention to provide IOL materials having extremely low coefficients of friction that are suitable for insertion through either lubricated, non-lubricated, or small-bore inserter barrels. Moreover, it is an object of the present invention that these low coefficient of friction IOL materials possess excellent biocompatibility and optical qualities.